The intra-aortic balloon pump is a cardiac assist device which operates on the counterpulsation principal. The purpose of the device is to temporarily assist the pumping function of a diseased or damaged heart. The device consists of a catheter with a long narrow balloon attached toward the distal end thereof for insertion through a femoral artery into the descending aorta. The proximal end of the catheter is connected to an external control system which provides timing and volume control for inflation and deflation of the balloon with helium gas. Following its initial acceptance circa 1970, the balloon catheter has become an important adjunct in the treatment of cardiac failure resulting from myocardial infarction or surgical intervention.
In operation, the intra-aortic balloon assists the blood pumping function of the heart with the timed inflation and deflation of the implanted balloon. The balloon is inflated as the heart completes its blood ejection cycle, resulting in an increased aortic blood pressure and systemic blood flow due to the displaced volume of the inflated balloon. Immediately prior to the next ejection cycle, the balloon is deflated and aortic blood pressure falls, reducing the workload on the heart during ejection.
Since its initial introduction, the design and construction of the balloon catheter has seen many changes and several manufacturers have developed products for the market. One dramatic change in the design and application of the balloon catheter eliminated the need for a surgical cut down to the femoral artery for insertion and made possible the insertion of the balloon through a percutaneous insertion device. For this insertion technique, the balloon is wrapped or folded into the smallest possible size or cross section and inserted into the femoral artery through a tubular system called a sheath/dilator assembly. The new percutaneous insertion technique allows insertion of the balloon in a broader range of medical situations and by a larger physician group. There is nearly universal acceptance of this important new technique. The new design and technique, however, is not totally problem free and balloon leaks in particular have been a point of concern among user physicians.
In operation, the intra-aortic balloon is inflated and deflated with helium gas during each cycle of the heart, representing many thousands of cycles during the average patient application. Rupture of the balloon membrane, causing gas leakage into the blood stream or blood leakage into the balloon system can have disastrous results and must be avoided. Balloon membrane reliability is a key element in all balloon designs and balloon rupture was in initial balloon designs an extremely rare occurrence. However, the recent incidence of balloon rupture has been reported by some users to be as high as five percent and has become a matter of significant concern. It is also recognized that this increased incidence of balloon rupture is connected with the change from surgically implanted balloons to percutaneously inserted balloons.
A review of balloon rupture going back to the beginnings of balloon pump use shows that balloon rupture, although rare, was seen in all balloon designs. When these early ruptured balloons were examined in the laboratory the rupture was determined to be the result of abrasion. The examination of recent balloons has shown the same characteristic abrasions and are believed to be the result of contact with calcified vascular lesions in the region of the renal arteries and bifurcation of the femoral arteries. The presence of calcified atherosclerosis in the aorta is well known and attributed to arterial stresses acting over long periods of time. It is also known that these vascular lesions occur first in the areas of highest stress, including the lower aorta around the origin of major vessels. Evidence of abrasion on intra-aortic balloons, removed from patients, is also found predominantly in the proximal portion of the balloon that had been positioned in the lower aorta.
The increased incidence of balloon rupture is thought to be the result of two conditions. First, balloons designed for percutaneous insertion have approximately half the wall or membrane thickness of the originals designed for surgical insertion. Second, the average age of the balloon pump patient population has increased and calcification in the lower aorta is more common and extensive.
The common intra-aortic balloon, when inflated, has a long cylindrical membrane shape with the maximum allowable length of the balloon being restricted by the need to avoid obstruction of the left subclavian artery and the renal arteries. The conservative establishment of this length is based on physiologic measurements in a broad patient population and any extension of the balloon length could result in obstruction of major arteries.
As the common shape of intra-aortic balloons currently marketed are cylindrical and are of constant diameter and wall thickness over the entire length, and since the length of the balloon is fixed, the displacement volume of the balloon is required for a 40 c.c. displacement, the most common balloon size, is approximately 0.57 inches. This diameter is close to the diameter of the aorta in the region of the renal arteries and brings the balloon wall within close proximity to the aorta wall. A protrusion of calcified atherosclerotic plaques can therefore come in nearly constant contact with the balloon during the inflation cycle.
A patent to R. T. Jones, U.S. Pat. No. 3,504,662, issued Apr. 7, 1970, very early in the development of intra-aortic balloon pumps and prior to development of the percutaneous insertion technique now practically universally accepted, broadly refers to varying the diameter of a compartmented membrane along the length thereof to provide a tapered configuration for the stated purpose of fitting within the aorta. The problem of abrasion-induced rupture was not recognized in Jones, and in fact did not appear until substantially later as a result of the development of the percutaneous insertion technique which required substantially thinner wall thicknesses for the membrane. This in turn led to an increase in the incidence of rupture and ultimately an extensive investigation into the causes thereof. The major factor in membrane rupture was found to be abrasion resulting from the presence of calcified atherosclerosis coupled with the thinner wall or membrane thickness required for percutaneous insertion.
Balloons of the type originally used in conjunction with femoral artery insertion by utilization of a surgical cut, because of the membrane thickness thereof, cannot be sufficiently wrapped or folded for percutaneous insertion. This would be the case even were a tapered balloon, as in Jones, used in that the membrane thickness at the wider diameter end portion of the balloon would, when folded or wrapped, present a substantial and unmanageable diametric bulk, even assuming the tapered end length is sufficiently collapsible. With regard to the thin membrane balloons, which incidentally are of only approximately one-half the membrane thickness of the earlier balloons as exemplified by Jones, there has been no successful resolution of the problem of abrasion and the rupturing resulting therefrom.